Quantitative measurement of specific biomolecules in complex biological samples is a key component of many molecular diagnostic tests. However, the challenge of making suitable measurements across many different analyte types in a manner that is compatible with the development of low cost, rapid-readout, hand-held, battery-operated point-of-care devices remains substantial.
For example, in the infectious disease field, there is an urgent need for rapid point-of-care devices that can provide information useful in the diagnosis of diseases such as tuberculosis (TB), malaria, HIV, etc. Amongst the many limitations of current potential point-of-care diagnostic tests in the TB field (and more generally) are:                Inadequate sensitivity of assay in terms of the concentration of analyte required for detection, leading to false negative results;        Lack of specificity of detecting reagents, leading to false positive results;        Poor temperature stability of detecting reagents, leading to false negative and false positive results.        
By way of illustration of such limitations, Dheda et al observed that existing antibody based Enzyme Linked Immuno Assays (ELISAs) for Mycobacterium tuberculosis (M. tb)-derived lipoarabinomannan (LAM), a complex glycolipid, work in urine with 99% specificity but with only 13% sensitivity, resulting in an unacceptably high false negative rate for use in the clinic.1
By comparison, the same test in sputum gave ˜86% sensitivity but only ˜15% specificity, resulting in an unacceptably high false positive rate. Dheda et al determined that the false positives in sputum are due to cross-reactivity of the anti-LAM antibodies used in the assay with LAM-like polymers produced by other microbes (both pathogenic and non-pathogenic) that co-habit the oral cavity.1 In contrast, the false-negative results are likely due to the combination of the intrinsic antibody-antigen affinity, combined with a low antigen concentration in the biological specimen and the intrinsic limit of detection of the ELISA method. The ability to rapidly and accurately measure the concentration of specific M.tb-derived compounds such as LAM in patient specimens whilst removing false negative results (by improving the limit of detection) and false positive results (by providing direct information on the identity of the analyte that has been captured) would therefore be a major advance in TB diagnostic tests, yet it is not practically possible today.
In a different example, a major challenge in the surgical field lies in administering the correct dose of anaesthetics to patients. It is well known that individual patients metabolise the majority of drug-like molecules (including anaesthetics) at widely varying rates due to inter-individual polymorphic variations present in for example the cytochrome P450 enzymes.2 As a consequence, the blood plasma concentrations of the active form of drugs such as anaesthetics (e.g. propofol) can vary widely between hyper- and null-metabolisers, leading in turn to variable responses to drug administration, the extreme results during surgery being either that a patient comes round during operation because the administered dose was too low for their genotype, or that the patient dies because the administered dose was too high for their genotype. In the absence of quantitative pharmacogenomic data on each individual patient that enables prior calculation of the exact optimal dose, the ability to rapidly and accurately measure and monitor the individual patient's blood plasma concentration of compounds such as propofol in real time in the operating theatre would therefore be a major advance in anaesthesiology, yet it is not practically possible.
A number of the shortcomings identified above in existing potential point-of-care diagnostic tests can be addressed through use of a novel Surface Enhanced Raman Scattering (SERS) assay platform, as described below.
Surface Enhanced Raman Scattering is a well known vibrational spectroscopy technique that has attracted considerable attention for its ultra sensitive, extremely specific and low limit of detection of biomolecules;3 it has been reported that, compared to traditional Raman spectroscopy, the ensemble averaged Raman signal in SERS increases 8-orders of magnitude, making it able in principle to detect single molecules.4 The SERS phenomenon utilises the intense localised evanescent wave (an electromagnetic field) that can be produced at metal surfaces and junctions by optical excitation of the surface plasmons to obtain a Raman spectrum or “signature” of surface adsorbed molecules. Classically, SERS measurements are made on individual pure compounds that are ‘Raman active’ and which are localised on an appropriate metal surface within the effective range of the evanescent wave. Typically a noble metal such as gold or silver is used as a SERS surface, but other transition metals such as copper iron, cobalt, nickel, palladium, and platinum can also be used.5 Since the propagation of an evanescent wave decays exponentially with distance from the from the boundary at which the wave was formed, SERS measurements are typically made on compounds localised within 20 nm of the metal surface,3,6 although SERS enhancement has been reported at distances up to 120 nm.7 Importantly, because of the direct relationship of the Raman shift of incident photons to the structure of the molecule under examination, the SERS technique is highly selective and each molecule has a distinct Raman signature that is also quantifiable. Thus in principle, SERS can be used to determine the identity of a compound (by comparing the measured SERS spectrum to a database of reference SERS spectra) as well as to measure its concentration.
Detection of biomolecules (including biomarkers) by SERS could thus potentially significantly improve both the sensitivity and specificity of diagnostic assays by providing quantitative information on the identity of the molecule being detected, whilst also providing lower limits of detection. However, when applied to complex mixtures of different molecules, overlapping SERS spectra derived from the different components of the mixture makes the task of identifying and quantifying individual components in the mixture essentially impossible without some prior separation or partitioning step; this consideration has limited the application of SERS to medical diagnostics to date.
A number of studies have shown that micro-fluidics combined with SERS can be used to detect trace explosives.8 It has also been reported that SERS can be used for various applications including detecting pollutants and DNA, whilst a SERS nano-biosensor has been designed that can accurately detect blood glucose at very low concentrations.9,10 
Some academic groups have attempted to enhance the detection capability of SERS by combining it with aptamers as a separation and enrichment matrix for specific molecules.11,12 Aptamers are oligonucleic acid or peptide molecules that bind to a specific target molecule and fold into 3D conformations in the presence of the target analytes.13 In particular, DNA aptamers are highly stable nucleic acid-based polymers that can bind in a high affinity and highly discriminatory manner to proteins, nucleic acids, carbohydrates, lipids and small molecules; their molecular recognition properties thus rival and possibly exceed those of antibodies, whilst probably being more compatible than antibodies with SERS due to their smaller physical size (DNA aptamers are typically ≦100 nt in length with a molecular weight ≦35 kDa). DNA aptamers are usually generated through use of in vitro selection methods and typically show greater thermo- and humidity-tolerance than antibodies because of their smaller size, the intrinsic stability of the phosphodiester linkage, and because they typically adopt a folded conformation reversibly in response to the presence of the cognate antigen.
Cho et al11 used an aptamer-based SERS sensor to detect thrombin. In their approach, a methylene blue-labelled anti-thrombin aptamer was first physically adsorbed to gold nano-particles; with the methylene blue-labelled aptamer in proximity to the gold surface, SERS of the methylene blue—a Raman-active dye—could occur. However, in the presence of thrombin, the anti-thrombin aptamer underwent a conformational change that weakened the physical association with the gold surface such that the aptamer-analyte complex (and hence the methylene blue label) diffused away from the surface and quenched the SERS signal. The resulting decrease in the methylene blue SERS signal was thus taken as an indirect indication of the binding of the aptamer to thrombin.11 
In a different approach, Huh & Erickson12 first labelled the protein vasopressin with the Raman-active dye FITC; when FITC-labelled vasopressin bound to an immobilised anti-vasopressin aptamer, the FITC label was brought into proximity of the surface, enabling the strong SERS signal of the FITC dye to be measured, thus giving indirect data on binding of vasoporessin to the aptamer.12
Notably, both the aptamer-SERS assays described above involve either the displacement of a Raman-labelled aptamer from a gold surface11 or the binding of a Raman-labelled protein to an immobilised aptamer12. Fundamentally both methods therefore monitor movements of the Raman label rather than directly monitoring the specific aptamer-ligand capture event itself. As such, those assays provide no information about the identity of the analyte bound by the aptamer, only that something has bound, and so do not differ fundamentally in information content from existing ELISA tests or other fluorescent detection techniques.
Neumann et al14 described the SERS-based detection of aptamer conformational changes induced by binding of the aptamer to target molecules such as proteins or organic ligands. In that work, Neumann et al demonstrated that the SERS spectrum of an unbound, thermally-denatured aptamer presented on a C6-alkyl thiol self-assembled monolayer (‘SAM’, ‘aptamer-SAM’) is reproducibly dominated by the adenine ring breathing mode of the aptamer, but noted that on binding of a specific ligand, the SERS spectrum of the aptamer-SAM became altered in an apparently poorly reproducible manner.14 Neumann et al thus aimed to deduce the binding of a target molecule to an immobilised aptamer by measuring the aptamer-SAM SERS spectrum of the unbound, thermally-denatured aptamer-SAM and then determining the apparent loss of reproducibility of the resultant aptamer-SAM SERS spectrum that occurs on ligand binding.14 Using circular dichroism spectroscopy, Neumann et al demonstrated for example that measurable conformational changes can be induced in an anti-cocaine aptamer-SAM by the specific target molecule cocaine, but also by the related but different molecules benzocaine and caffeine.14 As before therefore, the SERS method of Neumann et al provides no direct spectroscopic information on the identity of the aptamer-analyte complex itself; instead Neumann et al merely infer that something has bound to the aptamer-SAM (e.g. cocaine, benzocaine or caffeine in their example) and induced an apparently poorly reproducible change in the aptamer-SAM SERS spectrum.
There remains a need for detecting and measuring the amount of a given analyte in a complex biological sample that might, for instance, also include other molecules that can cross-react with the given analyte, giving rise to false positive data in other assays.